This invention relates generally to magnetic resonance imaging (MRI) systems and, more particularly, to radio-frequency (RF) coils in such systems.
Magnetic Resonance Imaging (MRI) utilizes hydrogen nuclear spins of the water molecules in the human body, which are polarized by a strong, uniform, static magnetic field generated by a magnet (typically denoted as B0—the main magnetic field in MRI physics). The magnetically polarized nuclear spins generate magnetic moments in the human body. The magnetic moments point or are aligned parallel to the direction of the main magnetic field B0 in a steady state and produce no useful information if they are not disturbed by any excitation.
The generation of nuclear magnetic resonance (NMR) signals for MRI data acquisition is accomplished by exciting the magnetic moments with a uniform radio-frequency (RF) magnetic field (typically referred to as the B1 field or the excitation field), for example, by applying a uniform RF magnetic field orthogonal to B0. This RF field is centered on the Larmor frequency of protons in the B0 field and causes the magnet moments to mutate their alignment away from B0 by some predetermined angle. The B1 field is produced in the imaging region of interest typically by an RF transmit coil that is driven by a computer-controlled RF transmitter with a RF power amplifier. During excitation, the nuclear spin system absorbs magnetic energy, and the magnetic moments precess around the direction of the main magnetic field. After excitation, the precessing magnetic moments will go through a process of free induction decay (FID), releasing their absorbed energy and returning to a steady state. During FID, NMR signals are detected by the use of a receive RF coil that is placed in the vicinity of the excited volume of a human body. The NMR signal is an induced electrical motive force (electrical voltage or current) in the receive RF coil that has been induced by the flux change over a period of time due to the relaxation of precessing magnetic moments of the human tissue. The receive RF coil can be either the transmit coil itself or an independent receive-only RF coil. The NMR signal is used for producing MR images by using additional pulsed magnetic gradient fields that are generated by gradient coils integrated inside the main magnet system. The gradient fields are used to spatially encode the signals and selectively excite a specific volume of the human body. There are usually three sets of gradient coils in a standard MRI system that generate magnetic fields in the same direction of the main magnetic field and vary linearly in the imaging volume.
In MRI, it is desirable for the excitation and reception to be spatially uniform in the imaging volume for better image uniformity. In a standard MRI system, the best excitation field homogeneity is usually obtained by using a whole-body volume RF coil for transmission. The whole-body transmit coil is the largest RF coil in the system. A large coil, however, produces lower signal-to-noise ratio (SNR or S/N) if it is also used for reception, mainly because of its greater distance from the signal-generating tissues being imaged. Because a high signal-to-noise ratio is the most desirable in MRI, special-purpose coils are used for reception to enhance the SNR ratio from the volume of interest.
In practice, a well-designed specialty RF coil has the following functional properties: high SNR ratio, good uniformity, high unloaded quality factor (Q) of the resonance circuit, and high ratio of the unloaded to loaded Q factors. In addition, the coil device must be mechanically designed to facilitate patient handling and comfort, and to provide a protective barrier between the patient and the RF electronics. Another way to increase the SNR is by quadrature reception. In this method, NMR signals are detected in two orthogonal directions that are in the transverse plane or perpendicular to the main magnetic field. The two signals are detected by two independent individual coils that cover the same volume of interest. With quadrature reception, the SNR can be increased, for example, by up to √2, over that of the individual linear coils.
To cover a large field-of-view, while maintaining the SNR characteristic of a small and conformal coil, linear surface coil arrays are used, for example, to image the entire human spine or to provide cervical, thoracic, lumbar (CTL) spine imaging. These linear surface coil arrays typically include an array of planar linear surface coil elements. However, these coil systems do not operate well for imaging deep tissues, such as, for example, the blood vessels in the lower abdomen, due to the sensitivity drop-off away from the coil surface.
To image the lower extremities, for example, lower human extremities, quadrature phased array coils have been utilized. One known quadrature phased array coil images the lower extremities by using two orthogonal linear coil arrays, for example, six planar loop coil elements placed in the horizontal plane and under the patient, and six planar loop coil elements placed in the vertical plane and in between the legs of the patient. Another known quadrature phased array coil images, for example, the blood vessels from the pelvis of patient and lower. This quadrature phased array coil includes two orthogonal linear coil arrays extending in the head-to-toe direction, for example, a planar array of loop coil elements positioned laterally and centrally located above a second array of butterfly coil elements. The loop coil elements, for example, are placed immediately under the patient and the butterfly coil elements are wrapped around the patient.
In MRI, gradient coils are routinely used to provide phase-encoding information to a sample to be imaged. To obtain an image, all of the data points in a so-called “k-space” (i.e., frequency space) must be collected. Some of the data points in the k-space are intentionally skipped and at the same time the intrinsic sensitivity information of the RF receive coils is used as the phase-encoding information for the skipped data points. This operation occurs simultaneously and, thus, is referred to as partially parallel imaging or partially parallel acquisition (PPA). By collecting multiple data points simultaneously, less time is required to acquire the same amount of data (e.g., image data) as compared to a conventional gradient-only phase-encoding approach. The time savings can be used to reduce total imaging time, in particular, for applications in which cardiac or respiratory motions in tissues being imaged become concerns, or to collect more data to achieve better resolution or SNR. PPA may be used to improved both temporal resolution (e.g., shorter acquisition time) and spatial resolution (e.g., collect more data points while the acquisition time remains unchanged).
Simultaneous Acquisition of Spatial Harmonics (SMASH) and Sensitivity Encoding (SENSE) are two methods of PPA. The two methods also may be combined to produce a hybrid method. SMASH utilizes parallel imaging by skipping phase encode lines that yield a decreasing field-of-view (FOV) in the phase-encoding direction and uses coils (e.g., coil arrays) together with reconstruction techniques to fill in the missing data points in the k-space. SENSE, on the other hand, utilizes a reduced FOV in the read direction, resulting in an aliased image that is then unfolded in the x-space (i.e., real space), while using the RF coil sensitivity information to obtain a true corresponding image. SENSE uses the phase and magnitude difference between signals from multiple coils to skip some of the phase encoding steps. By skipping some of the phase encoding steps, the image acquisition time is reduced, for example, by a reduction factor R, with the factor R equaling the number of independent coils/arrays. In SENSE imaging, the SNR is defined as: SNRSENSE=SNRFULL/{g√R}, where SNRFULL is the SNR that may be achieved when all the phase encoding steps are collected by a traditional gradient phase encoding scheme. SNRSENSE is optimized when the geometry factor g equals one. To obtain a “g” of one, traditional decoupling techniques such as overlapping nearest adjacent coil elements to null the mutual inductance between the coils may have to be relaxed.
The SENSE and SMASH or a hybrid approach have different RF coil design requirements than traditional coil designs. For example, in SMASH, the primary criterion for the coil array is that it is capable of generating sinusoids whose wavelengths are on the order of the FOV. Thus, the target FOV along the phase encoding direction for the array is determined in this manner. Conventional array designs can incorporate element and array dimensions that will give optimal SNR for the object of interest. Furthermore, conventional array designs may select almost any FOV so long as severe aliasing artifacts are not a concern. On the other hand, when using SMASH, the size of the array determines the approximate range of FOVs that can be used in the imaging test. This then determines the approximate element dimensions, assuming complete coverage of the FOV is desired, as in most cases.
SENSE is based upon the principle that the sensitivity of a RF receiver coil generally has a phase-encoding effect complementary to those achieved by linear field gradients. For SENSE imaging, the elements of a coil array may be smaller than for common multi-array imaging, resulting in a trade-off between basic noise and geometry factor. Further, when using SENSE, adjacent coil elements may not be overlapped for a net gain in SNR due to the improved geometry factor.
Thus, PPA design requirements are different than for traditional coil arrays based in part on the different operating characteristics. For example, SNR is intrinsically reduced when imaging time is reduced. Thus, to compensate for the SNR loss, the size of coil elements in PPA coil arrays tend to be smaller than the elements of a conventional array element.
Some known PPA coil arrays are based upon traditional RF coil design requirements using conventional coil design schemes and thus have the limitations of conventional systems (e.g., increased acquisition time). Further, some known PPA coil array designs include non-overlapping adjacent coil elements in an attempt to provide better definition in the individual phase and magnitude information associated with each RF coil used in the array, for example, to obtain information of B1 of a receive coil in SENSE applications. Without overlap, a net gain in SNR may be realized when using SENSE imaging, but the coupling between coils may also increase.
Thus, these known PPA coil arrays are not optimized for parallel imaging operation and limit the design options for such arrays. Therefore, the operation and control of an associated MRI system, operating, for example, using SENSE or SMASH is limited when using these known coil arrays.